In PET imaging, or positron emitter tomography, a radiopharmaceutical agent is administered, via injection, inhalation and/or ingestion, to a patient. The physical and bio-molecular properties of the agent then concentrate at specific locations in the human body. The actual spatial distribution, intensity of the point and/or region of accumulation, as well as the kinetics of the process from administration and capture to eventual elimination, all have clinical significance. During this process, the positron emitter attached to the radiopharmaceutical agent emits positrons according to the physical properties of the isotope, such as half-life, branching ratio, etc. Each positron interacts with an electron of the object, is annihilated and produces two gamma rays at 511 keV, which travel at substantially 180 degrees apart. The two gamma rays then cause a scintillation event at a scintillation crystal of the PET detector, which detects the gamma rays thereby. By detecting these two gamma rays, and drawing a line between their locations or “line-of-response,” the likely location of the original annihilation is determined. While this process only identifies one line of possible interaction, accumulating a large number of these lines, and through a tomographic reconstruction process, the original distribution is estimated with useful accuracy. In addition to the location of the two scintillation events, if accurate timing—within few hundred picoseconds—is available, time-of-flight calculations are also made in order to add more information regarding the likely position of the annihilation event along the line. Limitations in the timing resolution of a scanner determines the accuracy of the positioning along this line. Limitations in the determination of the location of the original scintillation events determines the ultimate spatial resolution of the scanner. A specific characteristic of the isotope (for example, energy of the positron) contributes (via positron range and co-linearity of the two gamma rays) to the determination of the spatial resolution for a specific radiopharmaceutical agent.
The above process is repeated for a large number of annihilation events. While every case needs to be analyzed to determine how many scintillation events are required to support the desired imaging tasks, conventionally a typical 100 cm long, FDG (fluoro-deoxyglucose) study accumulates about 100 millions counts or events. The time required to accumulate this number of counts is determined by the injected dose, as well as the sensitivity and counting capacity of the scanner.
PET imaging relies on the conversion of gamma rays into light through fast and bright scintillation crystals, generating the scintillation events referred to above. Time-of-Flight PET further requires sub-nanosecond timing resolution and resolutions of a few hundred picoseconds is also being contemplated. While it is complicated enough to tune and adjust two channels of scintillating crystal, photomultiplier tubes (PMT) and electronics, this complexity is only increased on a large arrays of crystals and sensors.
Modern PET systems support 500-600 ps timing resolutions. At this level, even small timing variation in the components are significant, and transit time is the most important variable in this equation. Transit time is the average time between when a photon strikes the photocathode of a PMT and when the corresponding current pulse is measured at the anode of the PMT. The variation of this quantity from one PMT to another causes the signals to reach the analysis circuitry at different times.
The need for an accurate transit time of the detection chain is often offset by internal or intrinsic ballistic differences between the crystal location relative to shortest and longest optical paths to the sensor. This is a complex theoretical estimation to perform, but measurements suggests that 25 to 40 ps is inherent timing variation related to optical path. Therefore, 25-40 ps accuracy in balancing the transit time of all the channels of the detector is a reasonable target. Any additional accuracy is useful, but has a marginal—if not negligible—effect on system performance.
Conventionally, several ways exist to control or add time delay to a PMT pulse in gamma ray detectors. Most methods involve active components, which degrade the frequency content of the signal and target unnecessary accuracy to the synchronization of all signals. Other conventional systems do not compensate for transit time variations between different PMT assemblies, and, therefore, suffer from degraded timing resolution. Moreover, the active circuitry approaches in conventional systems are costly, add to the complexity of the circuitry and, more importantly, degrade the quality and integrity of the signal.